Binary screen, system and method for single pulse dual energy radiography

ABSTRACT

A binary screen, system and method for performing dual energy radiographic imaging, using single pulse x-rays, are depicted. The binary screen includes a first phosphor which emits light of a first wavelength and a second phosphor which emits light of a second wavelength. The first and second phosphors are heterogeneously dispersed in the binary screen. The first phosphor is preferably Gd 2  O 2  S:Eu +3  and the second phosphor is preferably Y 2  O 2  S:Gd +3 . The second phosphor can also be Y 2  O 3  :Gd, Y 2  O 2  S:Pr or Y 2  O 2  S:Neodymium.

RELATED APPLICATIONS

The present application is a continuation-in-part application of Ser.No. 718,601, filed Jun. 20, 1991, now abandoned.

FIELD OF THE INVENTION

The present invention relates generally to the field of radiography andmore particularly, to dual energy radiography wherein a single pulse ofx-rays are utilized to generate multiple images.

BACKGROUND OF THE INVENTION

Dual-energy (DE) radiography is a well-known technique which allows thecalculation of either bone-only or soft tissue-only images. Thepotential and clinical effectiveness of DE radiography has been welldocumented. DE radiography requires the acquisition of two images inwhich the effective energy of the detected x-rays for each imagediffers. The desired calculation is achieved by subtracting one imagefrom the other. For a more detailed explanation of dual energyradiography, see W.R. Brody et al., A method for selective tissue andbone visualization using dual energy scanned projection radiography,Medical Physics, Vol. 8(3), May/June 1981, pps. 353-357. Presently thereare two practical acquisition scenarios for dual energy imaging.

In one category of DE image acquisition, two images are acquiredsequentially using a two-pulse method, wherein the kilovoltage and/orfiltration is changed between the two pulses of x-rays, thereby changingthe energy of the two images. Because of the problem of patient motion,the two images need to be acquired in rapid succession. This mandates anear real-time detector system such as an image intensifier (II) digitalvideo system, as well as an x-ray generator capable of rapid kVswitching. Van Lysel has recently reported on this type of dual energyacquisition for dual energy subtraction in cardiac angiography,following earlier studies by others. While useful for many angiographicapplications, II based imaging cannot meet either field of view orresolution requirements of general radiography.

Another category of dual energy acquisition utilizes a single x-raypulse and stacked detectors. See D.L. Ergun et al., Single-ExposureDual-Energy Computed Radiography: Improved Detection and Processing,Radiology, Medical Physics, Vol. 174, No. 1, 1990, pps. 243-249. Thepolyenergetic nature of the x-ray spectrum is exploited where a stack ofdetectors (with perhaps intervening filters) is used such that the firstdetector captures the lower energy image, and as the x-ray beam passesthough the first detector and intervening filtration it is hardenedbefore striking the second detector, which captures the higher energyimage.

The single pulse used in this technique is attractive because thepotential for motion artifacts is reduced. However, because thedetectors are necessarily stacked they need to be very thin and thusonly film/screen systems or stimulable phosphor systems can be used.Because stimulable phosphor systems to date are composed of only asingle phosphor, typically BaFBr:Eu, differences in the so-called k-edgeabsorption properties cannot be exploited. In order to digitize theimages from these stacked detector systems, they need to be separatedand in the process the spatial alignment between the detectors is lost.This requires the re-registration of the images on a case by case basis.

While filters with different k-edges can be placed in a stacked detectorbetween the fore and aft screens, the aft screen achieves energyseparation (relative to the fore screen) only through the removal ofx-ray photons from the x-ray beam, i.e, through filtration.Consequently, the high energy image derived from the aft screen willgenerally suffer from x-ray quantum noise effects, relative to the lowenergy image. Since the signal to noise ratio (SNR) in the subtractedimage is maximized when the number of absorbed x-ray quanta areapproximately equally distributed between the low energy and high energyimages, true optimization is hard to achieve using the stackedstimulable phosphor approach.

As indicated previously, the stacked stimulable phosphor cassette mustbe dismantled in order to read the latent images on each of the screensusing a scanning laser, requiring the images to be re-registered in thecomputer. Although correlation techniques have found to be a viableapproach to the re-registration, these techniques are numericallyintensive and could hinder practical clinical implementation of thetechnology. Furthermore, the scanning laser system employed to read outa stimulable phosphor screen is subject to time jitter noise similar tothose in video systems. Since CCD cameras employ an array of discretestationary detectors, time jitter noise does not exist for thesedevices. For more general background on the use of CCD cameras inradiography, see J.M. Heron et al., x-ray imaging with two-dimensionalcharge-coupled device (CCD) arrays, SPIE Vol. 486 Medical Imaging andInstrumentation '84, (1984), pps. 141-145.

Consequently, a need still exists for a dual energy radiography systemwhich does not suffer from x-ray quantum noise effects, which does notrequire dismantling in order to analyze images and which is capable ofreal time imaging.

SUMMARY OF THE INVENTION

The problems of prior DE radiographic devices are overcome and theadvantages of the invention are achieved in a binary screen, system andmethod for performing dual energy radiographic imaging, using singlepulse x-rays. The binary screen includes a first phosphor which emitslight of a first wavelength and a second phosphor which emits light of asecond wavelength. The first and second phosphors are heterogeneouslydispersed in the binary screen. The first phosphor is preferably Gd₂ O₂S:Eu⁺³ and the second phosphor is preferably Y₂ O₂ S:Gd⁺³. The secondphosphor can also be Y₂ O₃ :Gd, Y₂ O₂ S:Pr or Y₂ O₂ S:Neodymium.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention will be better understood, and its numerousobjects and advantages will become apparent to those skilled in the artby reference to the following detailed description of the invention whentaken in conjunction with the following drawings, in which:

FIG. 1 is a diagrammatic view of a dual energy radiographic system,constructed in accordance with the present invention;

FIG. 2 is a diagrammatic view of a single pulse x-ray and lightemissions relative to the binary screen contained in FIG. 1;

FIG. 3 is a graph of the relative intensity of Y₂ O₂ S:Gd and Gd₂ O₂S:Eu versus wavelength; and

FIG. 4 is a graph showing the energy dependence of the binary screen.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

A new and novel imaging display analysis device is shown in FIG. 1 andgenerally designated 10. A single pulse of x-rays is provided from asource 12 along axis A to and through a patient. The x-rays passingthrough the patient are incident upon binary screen 12. The compositionof binary screen 12 will be described in greater detail below. Lightemitted from binary screen 12 is detected by cameras 14 and 16, which inthe preferred embodiment are charge coupled device (CCD) cameras. Aswill be explained, camera 14 detects light of a first wavelength, whilecamera 16 detects light of a second wavelength. A shroud or cover 18 isprovided in order to prevent ambient light from interfering with cameras14 and 16. The image signals generated by cameras 14 and 16 are providedto processor 20. Processor 20 calculates a dual energy image in responseto the image signals. The dual energy image is displayed on display 22.

The dual energy acquisition system of the present invention utilizes asingle x-ray pulse but does not require stacked detectors. Thistechnique is specifically tailored for charge coupled device (CCD)camera acquisition, a technology which holds great promise in digitalradiography. The binary screen technique utilizes an intensifying screen12 with two rare-earth phosphors, each with a different elementalcomposition and hence different x-ray (k-edge related) energy absorptioncharacteristics. The two phosphors are thoroughly mixed together aspowders prior to being incorporated into the screen, resulting in auniform, heterogeneous distribution of the two phosphors in binaryscreen 12. Each phosphor is designed to emit a markedly differentwavelength of visible light than the other.

Two CCD cameras 14 and 16 are focused on screen 12. As shown in FIG. 1,camera 14 is spectrally sensitive to the emission of the low k-edgephosphor due to the use of optical filter 24. Camera 16 is sensitive toonly the emission wavelengths of the high k-edge phosphor due to the useof optical filter 26. The resulting parallel, two channel imaging systemshown in FIG. 2 can be used to acquire the high and the low energyimages simultaneously using one pulse of x-rays. There are no movingparts in this system, and the recycle time is dependent only on theread-out time of the CCD cameras. Furthermore, registration between theimages is achieved by mechanical alignment of the cameras and possibly adigital image warping technique to reduce spatial non-linearities, andonce calibrated there is no need for re-registration on an image byimage basis.

The binary screen approach has several distinct advantages overstimulable phosphor techniques. Because at present stimulable phosphorsemploy only one phosphor (Barium Fluoro-Halide), k-edge separationcannot be exploited. In other words, the relative concentration of thetwo phosphors in the binary screen can be adjusted to deliver comparablequantum absorption efficiency between the two phosphors at the optimalkV.

EXAMPLES

Before describing experimentation with the binary screen, binary screenabsorption will be described mathematically. The table below describesthe terms used:

    ______________________________________                                        Symbol       Description                                                      ______________________________________                                        I.sub.L (x,y)                                                                              Low energy image                                                 I.sub.H (x,y)                                                                              High energy image                                                f.sub.1      Weight fraction of phosphor 1                                    f.sub.2      Weight fraction of phosphor 2                                    μ.sub.1   Mass attenuation coefficient of                                               phosphor 1 (cm.sup.2 /gm)                                        μ.sub.2   Mass attenuation coefficient of                                               phosphor 2 (cm.sup.2 /gm)                                        t            Total mass thickness of the binary                                            screen (gm/cm.sup.2)                                             φ(E)     Photon spectrum emitted by the x-                                             ray tube (photons/cm.sup.2)                                      μ.sub.t   Attenuation coefficient of a pre-                                             patient filter                                                   t.sub.f      Filter thickness                                                 μ.sub.m   Attenuation coefficient of a mid-                                             screen filter                                                    t.sub.m      Mid-filter thickness                                             μ.sub.p   Attenuation coefficient of the                                                patient                                                          t.sub.p      Patient thickness                                                g.sub.1      Collective gain (and efficiency)                                              terms for phosphor 1                                             g.sub.2      Collective gain (and efficiency)                                              terms for phosphor 2                                             ______________________________________                                    

The low and high energy images generated using the binary screen aregiven by: ##EQU1##

These equations are analogous with those governing absorption by twodifferent attenuation processes, such as photoelectric and Comptonabsorption. If μ₁ (E)=μ₂ (E) in the above equations, no energyseparation would be possible. The photon spectrum incident upon thepatient should be tailored to the detector's energy sensitivities, byadjusting the kV of the x-ray generator and the composition andthickness of a pre-patient filter. Optimization along these lines willresult in a bimodal spectrum, with each mode tailored to the k-edges ofeach phosphor int eh binary screen. By comparison, the equationsgoverning the formation of the low and high energy images for stackedstimulable phosphor detector systems are: ##EQU2##

Contrary to the binary screen case, with photostimulable phosphordetectors there exists the condition that μ₁ (E)=μ₂ (E), because of thesingle phosphor used in photo-stimulable phosphor systems.

Computer simulations were used to examine properties of the binaryscreen system, and for comparison with photostimulable phosphor (PSP)systems. The simulations consisted primarily of numerical integrationover the energy spectrum, following the equations outlined above. Theattenuation coefficients for all elements were obtained from log-logpolynomial fit data tabulated by W.H. McMasters et al., Compilation ofx-ray cross sections, U.S. Dept. of Commerce, Springfield, VA 1969. Thex-ray spectra were generated using an algorithm, which is essentiallythe Birch and Marshall technique, see R. Birch et al., Computation ofbremsstrhlung x-ray spectra and comparison with spectra measured with aGe(Li) detector, Phys. Biol. and Med., Vol. 24, pps. 505 et seq. (1979).The computations were performed on a 486 PC (Northgate, Plymouth, MN). ADOS-extended C compiler (Itel C Code Builder) was used, which allowedtrue 32 bit code to utilize all 16 Megabytes of CPU RAM.

The dual energy subtraction algorithm used was adopted from D.L. Ergunet al., Single-exposure dual-energy computed radiography: Improveddetection and processing, Radiology, Vol. 125, pps. 243-245 (1977), andtakes the form:

    S=Ln(I.sub.H)-R Ln(I.sub.L)+k,                             (5)

where S is the signal. When the constants in Equation 5 are chosen suchthat: ##EQU3## Where the μ's (effective attenuation coefficients) arefor tissue, the signal S represents the tissue-subtracted image;

    S=(Rξ.sub.L -ξ.sub.H)t.sub.bone                      (7)

where ξ represents the attenuation coefficient for bone, and t_(bone) isthe thickness of bone.

It can be shown that the signal to noise ratio (SNR) for this procedureis given by: ##EQU4##

The computer simulations were performed to assess the optimal values forthe parameters of kV, pre-patient filter composition and thickness, andthe weight ratio of the first phosphor to the total phosphor thickness,w₁ /(w₁ +w₂). Since optimization in x-ray imaging implies the best imagequality for a given integral dose to the patient, a figure of merit(FOM) was chosen where: ##EQU5##

This FOM is independent of exposure. The dose to the patient wascalculated using parameterized energy absorption values from Monte Carlocalculations, which modeled the energy absorption in homogeneous waterphantoms of various thicknesses and of infinite lateral extent. Theabsorbed energy is the incident energy minus the sum of thebackscattered energy, forward scattered energy, and the transmittedprimary energy.

The binary screen simulations involved numerically integrating Equations1 and 2, and from these calculations the effective attenuationcoefficients and the number of absorbed photons (N₁ and N₂) weredetermined. The relative integral dose (i.e. absorbed energy in theobject) was also calculated. These values were then inserted intoEquation 9 via Equation 8 to assess the figure of merit for each set ofparameters. The PSP system simulations required numerical integrationusing Equations 3 and 4, which yielded the effective values of theattenuation coefficients and N₁ and N₂, which were also used to evaluatethe FOM via Equations 8 and 9. The PSP system simulations included twoother parameters, the mid-detector filtration atomic number andthickness. After preliminary studies, copper was chosen as themid-detector filter material (Z=29), and only the filter thickness wasincluded in the iterative optimization.

The prefilter thicknesses were constrained to keep the tube loadingfactor under 10. The binary screen detector shows a ridge of high FOMsfor prefilters of atomic numbers between about 58 and 68. A similarridge exists or the PSP system between Z=56 and Z=66. In both cases, theatomic number corresponding to the highest point on the FOM surfaceincreases gradually with increasing kV. Outside the relatively narrowrange of optimal prefilter atomic numbers, there is comparatively littleZ dependence of the FOM, for both binary and PSP systems. In contrast,the kV dependence of the FOM surfaces is relatively strong. The optimumvalues for all parameters corresponding to the best FOM surfaces isrelatively strong. The optimum values for all parameters correspondingto the best FOM for each object thickness are given in Table 1 for thebinary screen and in Table 2 for the PSP system.

                                      TABLE 1                                     __________________________________________________________________________    Binary Screen Parameters at Optimal FOM                                       100 mg/cm.sup.2 total screen thickness                                                Prefilter                                                             Object  Atomic                                                                             Prefilter                                                                           Phosphor E.sub.1                                                                           E.sub.2                                                                           ΔE                                  Thickness                                                                           kV                                                                              Number                                                                             Thickness.sup.1                                                                     Ratio.sup.2                                                                        FOM (keV)                                                                             (keV)                                                                             (keV)                                     __________________________________________________________________________    10 cm 64                                                                              58   200   90%  27,188                                                                            39.7                                                                              46.2                                                                              6.5                                                    mg/cm.sup.2                                                      20 cm 62                                                                              58   200   85%  640.8                                                                             41.8                                                                              48.5                                                                              6.7                                                    mg/cm.sup.2                                                      30 cm 64                                                                              60   200   80%  26.65                                                                             45.1                                                                              51.6                                                                              6.5                                                    mg/cm.sup.2                                                      __________________________________________________________________________     .sup.1 The prefilter thickness was constrained by requiring that the tube     loading factor be under 10.                                                   .sup.2 The phosphor ratio refers to the weight ratio of the Y.sub.2           O.sub.2 S phosphor to the total phosphor in the screen, w.sub.1 (w.sub.1      +w.sub.2).                                                               

                                      TABLE 2                                     __________________________________________________________________________    Photostimulable Phosphor System Parameters at Optimal FOM                     100 mg/cm.sup.2 total screen thickness                                        Copper midfilter material                                                             Prefilter                                                             Object  Atomic                                                                             Prefilter                                                                           Phosphor                                                                           Midfilter E.sub.1                                                                           E.sub.2                                                                           ΔE                            Thickness                                                                           kV                                                                              Number                                                                             Thickness.sup.1                                                                     Ratio.sup.2                                                                        Thickness.sup.3                                                                     FOM (keV)                                                                             (keV)                                                                             (keV)                               __________________________________________________________________________    10 cm 62                                                                              58   200   65   300   22,080                                                                            40.9                                                                              47.4                                                                              6.5                                              mg/cm.sup.2                                                                              mg/cm.sup.2                                           20 cm 68                                                                              60   200   60   300   502.5                                                                             46.5                                                                              53.2                                                                              6.7                                              mg/cm.sup.2                                                                              mg/cm.sup.2                                           30 cm 70                                                                              62   200   60   300   21.39                                                                             50.1                                                                              55.8                                                                              5.6                                              mg/cm.sup.2                                                                              mg/cm.sup.2                                           __________________________________________________________________________     .sup.1 The prefilter thickness was constrained by requiring that the tube     loading factor be under 10.                                                   .sup.2 The phosphor ratio refers to the weight ratio of the front screen      phosphor to the total phosphor in the screen, w.sub.1 (w.sub.1 +w.sub.2).     .sup.3 The optimal midfilter thickness was searched between 0-300             mg/cm.sup.2 .                                                            

The optimal spectra for the binary and PSP systems for a 20 cm objectthickness were determined. The incident spectrum is noticed at the 40keV k-edge of the cerium filtration, which produced a bimodal spectrum.In principle, the low energy lobe is tailored to the yttrium k-edge, andthe high energy lobe is matched of the gadolinium k-edge. Once theincident spectrum passed through the patient (i.e. 20 cm of water), thetransmitted spectrum showed a more even distribution in the number ofphotons between the two lobes of the bimodal distribution, due togreater attenuation at lower photon energies. Whereas the absorptionspectrum for the Y₂ O₂ S phosphor component in the binary screenessentially mirrored the spectral distribution incident upon it, theabsorption spectrum in the Gd₂ O₂ S phosphor reflected the increasedabsorption above the 50 keV k-edge of Gadolinium.

In order to make comparisons between theory and experiment, experimentalx-ray spectra were characterized using the equivalent spectrumtechnique. In this technique, the experimentally measured aluminumattenuation is compared iteratively with aluminum attenuation valuesderived from computer-generated spectra (based on measured spectra), andthe computer-generated spectrum which best matches the measured aluminumattenuation properties represents the characterized or so-calledequivalent spectrum.

Binary screens composed of Y₂ O₂ S:Gd⁺³ and Gd₂ O₂ S:Eu⁺³ weremanufactured for this project. The optical emission spectra weredetermined and the Y₂ O₂ S:Gd emission was primarily at 514 nm, (green)and the Gd₂ O₂ S:Eu emission was primarily at 624 nm (red). KodakWratten optical filters were used to isolate the emission wavelengths.

The manufactured binary screens included a support and the mixed binarycomposition applied in a dispersed state in a binder. The support can beformed from a film such as polyethylene terephthalate (PET) or fromother materials such as paper, cardboard, glass and metal. The screen isprepared by dispersing the phosphor particles in a solution of thebinder and coated on the support. The coating application can beperformed by any known technique such as doctor blade coating, sprayingor dip coating. Suitable binders which are transparent to x-rayradiation and emitted radiation are organic polymers such as acrylates,polyurethanes and polycarbonates. Phosphor coverage on the screen can bein the range of 30 to 150 mg/cm². It is also noted that a lightreflecting layer such as titanium dioxide may be provided between thephosphor containing layer and the support to enhance the output oflight. It is preferred that the support onto which the phosphor iscoated may be black or light absorbing in order to increase imagesharpness. Colored dyes and pigments can also be added to achieve thisresult. Other intensifying screen techniques can be used with theinvention such as overcoating and the use of subbing layers.

The linearity of the photodetectors was determined by comparing theirresponse with that of a radiometer (EG&G Gamma Scientific, San Diego,CA), when placed in a light field generated from an adjustable DChalogen source filtered by a 600 nm band-pass filter (Oriel). Thebandpass filter resulted in a pseudo-monoenergetic light source. Thebandpass filter was necessary to assure that the linearity curve was notdistorted by the minor differences in the spectral responses between thephotodiodes and the radiometer, since the spectral output of the lightsource does change with intensity (i.e., applied voltage). The signalvoltage from the photodiodes was digitized using a computer-mounted 12bit analog to digital converter (ADC) (Model 2801-A, Data Translation,Marlboro, MA). 512 samples were acquired with a sampling period of 50microseconds and averaged for the linearity study. The photodiodelinearity was determined over 4 decades of optical exposure. Linearregression analysis between the two photodiode responses and theradiometer yielded correlation coefficients of 0.99993 and 0.99998.Linear regression analysis on log-log data indicated γ of 1.029 and1.039 for the two photodiodes. The photodiodes exhibited excellentlinearity, and thus no correction techniques were necessary tocompensate for the characteristic curve of these photodetectors.

It should be noted that the linearity of the photodetectors was found tobe crucial in demonstrating the energy dependence of the dual detectorsystem. Because the intensity of light striking each of thephotodetectors fluctuates significantly as a function of kV,nonlinearities in the detector response will yield artifactual energyresponse curves. The photodetector arrangement described, was chosenafter prior attempts with phototransistors (which lacked enoughsensitivity) and different photodiodes (which proved too noisy).

Prior to the development of the Y₂ O₂ S:Pr/Gd₂ O₂ S:Eu screen describedabove, Y₂ O₃ :Gd/Gd₂ O₂ S:Eu binary screens were manufactured. The Y₂ O₃:Gd emits at 315 nm (ultraviolet), and therefore has better spectralseparation from the 624 nm Gd₂ O₂ S:Eu than the 514 nm Y₂ O₂ S:Pr.However, after extensive testing it is believed that the Gd₂ O₂ S:Euphosphor absorbed much of the UV emission of the Y₂ O₃ :Gd. Thedifficulties of detecting the UV emission were exacerbated by therelative insensitivity of photodetectors to UV and their largesensitivity to the red emission.

To measure the energy dependency of the binary screen as a function ofkilovoltage, two photodiodes were sampled simultaneously (multiplexedacquisition at 20 kHz) during an x-ray exposure. Each photodiode wasoptically filtered to select for either the green emission of the Y₂ O₂S:Gd phosphor or the red emission of the Gd₂ O₂ S:Eu phosphor. Thesignal acquisition was synchronized to the x-ray pulse by triggering offof the signal derived from a phototransistors coupled to a small (2 mm×2mm) piece of x-ray screen, mounted in the x-ray collimator. A singlephase, full wave rectified x-ray generator (Siemens Heliphos), coupledto a Machlett Dynamax dual focus (1.0/2.0 mm) was employed as the x-raysource. To avoid exposing the photodiodes to the x-ray beam, a frontsurfaced mirror was used to reflect the optical radiation to thephotodetectors, which were protected from the x-ray beam by 2 cm of Pb.

The single phase x-ray source made data analysis more complicated. Aftereach x-ray exposure was acquired, the 512 element waveform for eachphotodiode was computer processed, the peaks in the waveform wereselected, and the voltages at the peaks were averaged to yield thephotodiode output value. This peak-detect regime was used to simulatethe effects of x-ray generators with less voltage ripple, e.g. threephase and inverter types.

The energy dependence of the binary screen as determined from thephotodiode experiment is shown in FIG. 4. The ratio of the measuredintensities illustrates the energy dependence, and this is related tothe logarithmic difference signal used in the actual dual energy imagecalculations. As expected, the red emission from the Gd₂ O₂ S:Euincreases with increasing incident beam energy. The experiment was alsoperformed with no optical filtration, and this resulted in thehorizontal line of data on the FIG. 4, illustrating no energy dependencewhen the optical filtration is removed.

The isometric plots illustrating the functional dependence of the FOMson kV and filter composition are quite similar between the binary screensystem and the PSP systems. The most striking feature of these resultsis the sharp FOM ridges within a very narrow range of prefilter atomicnumbers. The range of Z's is remarkably similar between the two detectorsystems, considering that the k-edges of the detectors differsubstantially. The binary screen simulated here consisted of yttrium andgadolinium, based phosphors, with k-edges of 17 keV and 50 keV,respectively. The barium k-edge of the (BaFBr) PSP screens is at 37 keV.The FOMs are composed of both a SNR component and an integral dosecomponent. It was determined that when graphed, sharp FOM ridges can beattributed to an improvement in the SNR, and not a reduction in theintegral does.

The optimal kVs (for both detector systems) are low, compared toconventional x-ray imaging kVs, for thick objects. This is not reallysurprising, since the fidelity of dual energy images relies ondifferential energy dependence between the attenuation coefficients ofcalcium and tissue. There is a tendency in the dual energy imagingliterature to equate the energy separation between the low and the highimages (ΔE) as a loose measure of subtraction quality. In fact, a largeΔE at high energies is less valuable than a smaller ΔE at lowerenergies, where the attenuation coefficients have large functionaldependencies on energy. The signal amplitude was calculated as ##EQU6##The signal resulting from a ΔE of 5 keV was higher at 40 keV than thatresulting from a ΔE of 20 keV at 57 keV. Although the optimal dualenergy subtraction signal region is below 20 keV, this increased whenthe effects of noise and dose are included.

It will be noted that if Y₂ O₂ S:Pr is used it is further preferred tocombine this phosphor with the Gd₂ O₂ S:Eu phosphor in equal, i.e.50/50, proportions. In other words, the relative concentration of thetwo phosphors in the binary screen can be adjusted to deliver comparablequantum absorption efficiency between the two phosphors at the optimalkV. The optimal distribution of photons between the two phosphors is afunction of the weighting coefficient between the high and low energyimages. See R in Equation 5. When R=1 the optimal weighting coefficientis 0.5.

In sum, the present invention introduces the use of binary screen 12,enabling the use of dual CCD cameras in dual energy imaging. Twophosphors, Y₂ O₂ S:Gd phosphor and Gd₂ O₂ S:Eu phosphor are combinedinto a single, binary screen. The difference in k-edges results inenergy dependent x-ray absorption by each of the phosphors. Eachphosphor is designed to emit a different wavelength of visible light,514 nm (Green) and 624 nm (Red), respectively. The screen is opticallycoupled to dual CCD cameras 14 and 16, which are optically filtered suchthat one camera is sensitive to only the emission of the low k-edgephosphor and the other camera is sensitive to the emission spectra ofthe high k-edge phosphor. A dual channel system utilizing no movingparts and capable of simultaneous dual energy acquisition results.

While the invention has been described and illustrated with reference tospecific embodiments, those skilled in the art will recognize thatmodification and variations may be made without departing from theprinciples of the invention as described herein above and set forth inthe following claims.

What is claimed is:
 1. A binary screen or use in dual energyradiography, comprising;a first rare earth phosphor, wherein said firstx-ray phosphor emits light of a first wavelength, said first phosphorcomprising Gd₂ O₂ S:Eu; and a second rare earth phosphor, wherein saidsecond phosphor emits light of a second wavelength, said second phosphorconsisting of Y₂ O₃ :Ge, Y₂ O₂ S:Pr or Y₂ O₂ S:Neodymium wherebyspectral separation between said first and second phosphors is achieved.2. The screen of claim 1, further comprising a binder, wherein saidfirst and second phosphors are heterogeneously dispersed in said binder.3. The screen of claim 2, further comprising a support, wherein saidbinder containing said first and second phosphors is applied to saidsupport.